The present invention relates generally to diagnostic imaging and, more particularly, to a method and apparatus of detecting materials with K-edge absorption in an effective energy region covered by an incident x-ray spectrum.
Exemplary diagnostics devices comprise x-ray systems, magnetic resonance (MR) systems, ultrasound systems, computed tomography (CT) systems, positron emission tomography (PET) systems, and other types of imaging systems. Typically, in CT imaging systems, an x-ray source emits a fan-shaped beam toward a subject or object, such as a patient or a piece of luggage. Hereinafter, the terms “subject” and “object” shall include anything capable of being imaged. The beam, after being attenuated by the subject, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the x-ray beam by the subject. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element. The electrical signals are transmitted to a data processing system for analysis which ultimately produces an image.
Generally, the x-ray source and the detector array are rotated about the gantry opening within an imaging plane and around the subject. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal point. X-ray detectors typically include a collimator for collimating x-ray beams received at the detector, a scintillator for converting x-rays to light energy adjacent the collimator, and photodiodes for receiving the light energy from the adjacent scintillator and producing electrical signals therefrom.
Typically, each scintillator of a scintillator array converts x-rays to light energy. Each scintillator discharges light energy to a photodiode adjacent thereto. Each photodiode detects the light energy and generates a corresponding electrical signal. The outputs of the photodiodes are then transmitted to the data processing system for image reconstruction.
An exemplary CT imaging system comprises an energy discriminating (ED) and/or multi energy (ME) CT imaging system that may be referred to as an EDCT and/or MECT imaging system. The EDCT and/or MECT imaging system in an example is configured to be responsive to different x-ray spectra. For example, a conventional third generation CT system acquires projections sequentially at different x-ray tube potentials. Two scans in an example are acquired either back to back or interleaved in which the tube operates at 80 kVp and 160 kVp potentials. Special filters in an example are placed between the x-ray source and the detector such that different detector rows collect projections of different x-ray energy spectra. The special filters that shape the x-ray spectrum in an example can be used for two scans that are acquired either back to back or interleaved. Energy sensitive detectors in an example are used such that each x-ray photon reaching the detector is recorded with its photon energy.
Exemplary ways to obtain the measurements comprise: (1) scan with two distinctive energy spectra, (2) detect photon energy according to energy deposition in the detector, and (3) photon counting. EDCT/MECT provides energy discrimination and material characterization. For example, in the absence of object scatter, the system derives the behavior at any other energy based on the signal from two regions of photon energy in the spectrum: the low-energy and the high-energy portions of the incident x-ray spectrum. In an exemplary energy region of medical CT, two physical processes dominate the x-ray attenuation: (1) Compton scatter and the (2) photoelectric effect. The detected signals from two energy regions provide sufficient information to resolve the energy dependence of the material being imaged. Furthermore, detected signals from the two energy regions provide sufficient information to determine the relative composition of an object composed of two materials.
The conventional basis material decomposition (BMD) algorithm is based on the concept that in the energy region for medical CT, the x-ray attenuation of any given material can be represented by a proper density mix of two other materials, referred to as the basis materials. Based on the projections acquired at the two incident x-ray spectra, the BMD algorithm computes two sets of new projections, corresponding to two new CT images that each represents the equivalent density of one of the basis materials. Since a material density is independent of x-ray photon energy, these images are approximately free of beam-hardening artifacts. An operator can choose the basis material to target a certain material of interest, for example, to enhance the image contrast.
Medical CT images can be enhanced in certain applications by use of contrast agents. A contrast agent is injected and images can be taken below and above the K-edge absorption energy of the contrast agent to further the contrast agent. For example, the two images are logarithmically subtracted and show the details of the structure of those volumes containing the contrast agent.
A K-edge indicates a sudden increase in the attenuation coefficient of photons occurring at a photon energy just above the binding energy of the K shell electron of the atoms interacting with the photons. The sudden increase in attenuation is due to photoelectric absorption of the photons. For this interaction to occur, the photons have more energy than the binding energy of the K shell electrons. A photon having an energy just above the binding energy of the electron is therefore more likely to be absorbed than a photon having an energy just below this binding energy. A general term for the phenomenon is absorption edge.
Systems with K-edge contrast materials and/or agents do not fit into the conventional BMD model. The conventional BMD is directed to non K-edge materials. In addition, the conventional BMD cannot account for the K-edge effect of high Z or high atomic number materials such as iodine (I), barium (Ba), tungsten (W), gadolinium (Gd), and xenon (Xe) if the K-edge of the material lies in the active energy region of the incident x-ray spectrum. A design for resolving K-edge contrast agents has employed monochromatic x-ray beams with which the K-edge material can be resolved by imaging the object at photon energies slightly below and slightly above the K-edge, but prevents integration of monochromatic sources with sufficient x-ray flux into a rotating gantry and so limits the application from use as a practical monochromatic x-ray source in medical CT scanners. An exemplary K-edge material comprises a K-edge within an x-ray spectrum employed for a given, selected, and/or particular application. An exemplary non K-edge material may comprise no K-edge, or may comprise a K-edge that is outside the x-ray spectrum for such an application. For example, iodine comprises a K-edge at approximately 33.2 keV. Iodine does comprise a K-edge material in an exemplary low kVp system where the x-ray spectrum covers approximately 20 keV to approximately 50 keV. In another example, iodine would not be considered a K-edge material in a system where the x-ray spectrum starts from approximately 40 keV.
Developments in biotechnology show promise for contrast agents that target specific organs and/or diseases. These contrast agents can be designed to have high-Z elements with a K-edge above 50 keV. With this K-edge in the middle of the x-ray energy spectrum, the conventional BMD cannot account for discontinuity in the attenuation and fails to provide accurate results. Therefore, it would be desirable to design an apparatus and method to reduce a number of scans needed in diagnostic imaging to accurately detect a density of one or more K-edge contrast agents.